Drug release stent coating process

ABSTRACT

A method of coating implantable open lattice metallic stent prosthesis is disclosed which includes sequentially applying a plurality of relatively thin outer layers of a coating composition comprising a solvent mixture of uncured polymeric silicone material and crosslinker and finely divided biologically active species, possibly of controlled average particle size, to form a coating on each stent surface. The coatings are cured in situ and the coated, cured prosthesis are sterilized in a step that includes preferred pretreatment with argon gas plasma and exposure to gamma radiation electron beam, ethylene oxide, steam.

BACKGROUND OF THE INVENTION

I. Cross-Reference to Related Applications

The present application is a Continuation-In-Part of copendingapplication Ser. No. 08/526,273, abandoned, filed Sep. 11, 1995, and aContinuation-In-Part of copending application Ser. No. 08/424,884,abandoned filed Apr. 19, 1995, all portions of the parent applicationsnot contained in this application being deemed incorporated by referencefor any purpose. Cross-reference is also made to application Ser. No.08/663,518, entitled "DRUG RELEASE STENT COATING AND PROCESS", filed ofeven date and of common inventorship and assignee, that is also aContinuation-In-Part of both above-referenced patent applications. Anyportion of that application that is not contained herein is also deemedto be incorporated by reference for any purpose.

II. Field of the Invention

The present invention relates generally to therapeutic expandable stentprosthesis for implantation in body lumens, e.g., vascular implantationand, more particularly, to a process for providing biostable elastomericcoatings on such stents which incorporate biologically active specieshaving controlled release characteristics directly in the coatingstructure.

II. Related Art

In surgical or other related invasive medicinal procedures, theinsertion and expansion of stent devices in blood vessels, urinarytracts or other difficult to access places for the purpose of preventingrestenosis, providing vessel or lumen wall support or reinforcement andfor other therapeutic or restorative functions has become a common formof long-term treatment. Typically, such prosthesis are applied to alocation of interest utilizing a vascular catheter, or similartransluminal device, to carry the stent to the location of interestwhere it is thereafter released to expand or be expanded in situ. Thesedevices are generally designed as permanent implants which may becomeincorporated in the vascular or other tissue which they contact atimplantation.

One type of self-expanding stent has a flexible tubular body formed ofseveral individual flexible thread elements each of which extends in ahelix configuration with the centerline of the body serving as a commonaxis. The elements are wound in a common direction, but are displacedaxially relative to each other and meet, under crossing a like number ofelements also so axially displaced, but having the opposite direction ofwinding. This configuration provides a resilient braided tubularstructure which assumes stable dimensions upon relaxation. Axial tensionproduces elongation and corresponding diameter contraction that allowsthe stent to be mounted on a catheter device and conveyed through thevascular system as a narrow elongated device. Once tension is relaxed insitu, the device at least substantially reverts to its original shape.Prosthesis of the class including a braided flexible tubular body areillustrated and described in U.S. Pat. Nos. 4,655,771 and 4,954,126 toWallsten and 5,061,275 to Wallsten et al.

Implanted stents have also been used to carry medicinal agents, such asthrombolytic agents. U.S. Pat. No. 5,163,952 to Froix discloses athermal memoried expanding plastic stent device which can be formulatedto carry a medicinal agent by utilizing the material of the stent itselfas an inert polymeric drug carrier. Pinchuk, in U.S. Pat. No. 5,092,877,discloses a stent of a polymeric material which may be employed with acoating associated with the delivery of drugs. Other patents which aredirected to devices of the class utilizing bio-degradable orbio-sorbable polymers include Tang et al, U.S. Pat. No. 4,916,193, andMacGregor, U.S. Pat. No. 4,994,071. Sahatjian in U.S. Pat. No.5,304,121, discloses a coating applied to a stent consisting of ahydrogel polymer and a preselected drug; possible drugs include cellgrowth inhibitors and heparin. A further method of making a coatedintravascular stent carrying a therapeutic material in which a polymercoating is dissolved in a solvent and the therapeutic material dispersedin the solvent and the solvent thereafter evaporated is described inBerg et al, U.S. Pat. No. 5,464,650, issued Nov. 5, 1995 andcorresponding to European patent application 0 623 354 A1 published 09Nov. 1994.

An article by Michael N. Helmus (a co-inventor of the present invention)entitled "Medical Device Design--A Systems Approach Central VenousCatheters", 22nd International Society for the Advancement of Materialand Process Engineering Technical Conference (1990) relates topolymer/drug/membrane systems for releasing heparin. Those polymer/drug/membrane systems require two distinct layers to function.

The above cross-referenced grandparent application supplies an approachthat provides long-term drug release, i.e., over a period of days oreven months, incorporated in a controlled-release system. The parentapplication and present invention provide a process for coating suchstents including techniques that enable the initial burst effect of drugelation to be controlled and the drug release kinetic profile associatedwith long-term therapeutic effect to be modified.

Metal stents of like thickness and weave generally have bettermechanical properties than polymeric stents. Metallic vascular stentsbraided of even relatively fine metal filament can provide a largeamount of strength to resist inwardly directed circumferential pressurein blood vessels. In order for a polymer material to provide comparablestrength characteristics, a much thicker-walled structure or heavier,denser filament weave is required. This, in turn, reduces thecross-sectional area available for flow through the stent and/or reducesthe relative amount of open space available in the structure. Inaddition, when applicable, it is usually more difficult to load anddeliver polymeric stents using vascular catheter delivery systems.

It will be noted, however, that while certain types of stents such asbraided metal stents may be superior to others for some applications,the process of the present invention is not limited in that respect andmay be used to coat a wide variety of devices. The present inventionalso applies, for example, to the class of stents that are notself-expanding including those which can be expanded, for instance, witha balloon. Polymeric stents, of all kinds can be coated using theprocess. Thus, regardless of particular detailed embodiments the use ofthe invention is not considered or intended to be limited with respecteither to stent design or materials of construction. Further, thepresent invention may be utilized with other types of implantprostheses.

Accordingly, it is a primary object of the present invention to providea coating process for coating a stent to be used as a deployed stentprosthesis, the coating being capable of long-term delivery ofbiologically active materials.

Another object of the invention is to provide a process for coating astent prosthesis using a biostable hydrophobic elastomer in whichbiologically active species are incorporated within a cured coating.

Still another object of the present invention is to provide amulti-layer coating in which the percentage of active material can varyfrom layer to layer.

A further object of the present invention is to control or modifyaspects of the timed or time variable drug delivery from a stent coatingby controlling average particle size in the biologically active species.

Other objects and advantages of the present invention will becomeapparent to those skilled in the art upon familiarization with thespecification and appended claims.

SUMMARY OF THE INVENTION

The present invention provides processes for producing a relatively thinlayer of biostable elastomeric material in which an amount ofbiologically active material is dispersed as a coating on the surfacesof a deployable stent prosthesis. The preferred stent to be coated is aself-expanding, open-ended tubular stent prosthesis. Although othermaterials, including polymer materials, can be used, in the preferredembodiment, the tubular body is formed of an open braid of fine singleor polyfilament metal wire which flexes without collapsing and readilyaxially deforms to an elongate shape for transluminal insertion via avascular catheter. The stent resiliently attempts to resumepredetermined stable dimensions upon relaxation in situ.

The coating is preferably applied as a mixture, solution or suspensionof polymeric material and finely divided biologically active speciesdispersed in an organic vehicle or a solution or partial solution ofsuch species in a solvent or vehicle for the polymer and/or biologicallyactive species. For the purpose of this application, the term "finelydivided" means any type or size of included material from dissolvedmolecules through suspensions, colloids and particulate mixtures. Theactive material is dispersed in a carrier material which may be thepolymer, a solvent, or both. The coating is preferably applied as aplurality of relatively thin layers sequentially applied in relativelyrapid sequence and is preferably applied with the stent in a radiallyexpanded state. In some applications the coating may further becharacterized as a composite initial tie coat or undercoat and acomposite topcoat. The coating thickness ratio of the topcoat to theundercoat may vary with the desired effect and/or the elution system.Typically these are of different formulations.

The coating may be applied by dipping or spraying using evaporativesolvent materials of relatively high vapor pressure to produce thedesired viscosity and quickly establish coating layer thicknesses. Thepreferred process is predicated on reciprocally spray coating a rotatingradially expanded stent employing an air brush device. The coatingprocess enables the material to adherently conform to and cover theentire surface of the filaments of the open structure of the stent butin a manner such that the open lattice nature of the structure of thebraid or other pattern is preserved in the coated device.

The coating is exposed to room temperature ventilation for apredetermined time (possibly one hour or more) for solvent vehicleevaporation. Thereafter the polymeric precuser material is cured at roomtemperature or elevated temperatures or the solvent evaporated away fromthe dissolved polymer as the case may be curing is defined as theprocess of converting the elastomeric or polymeric material into thefinished or useful state by the application of heat and/or chemicalagents which include physical-chemical charges. Where, for example,polyurethane thermoplastic elastomers are used, solvent evaporation canoccur at room temperature rendering the polymeric material useful forcontrolled drug release without further curing. Non-limiting examples ofcuring according to this definition include the application of heatand/or chemical agents and the evaporation of solvent which may inducephysical and/or chemical changes.

The ventilation time and temperature for cure are determined by theparticular polymer involved and particular drugs used. For example,silicone or polysiloxane materials (such as polydimethylsiloxane) havebeen used successfully. These materials are applied as pre-polymer inthe coating composition and must thereafter be cured. The preferredspecies have a relatively low cure temperatures and are known as a roomtemperature vulcanizable (RTV) materials. Some polydimethylsiloxanematerials can be cured, for example, by exposure to air at about 90° C.for a period of time such as 16 hours. A curing step may be implementedboth after application of a certain number of lower undercoat layers andthe topcoat layers or a single curing step used after coating iscompleted.

The coated stents may thereafter be subjected to a postcuresterilization process which includes an inert gas plasma treatment, andthen exposure to gamma radiation, electron beam, ethylene oxide (ETO) orsteam sterilization may also be employed.

In the plasma treatment, unconstrained coated stents are placed in areactor chamber and the system is purged with nitrogen and a vacuumapplied to about 20-50 mTorr. Thereafter, inert gas (argon, helium ormixture of them) is admitted to the reaction chamber for the plasmatreatment. A highly preferred method of operation consists of usingargon gas, operating at a power range from 200 to 400 watts, a flow rateof 150-650 standard ml per minute, which is equivalent to about 100-450mTorr, and an exposure time from 30 seconds to about 5 minutes. Thestents can be removed immediately after the plasma treatment or remainin the argon atmosphere for an additional period of time, typically fiveminutes.

After the argon plasma pretreatment, the coated and cured stents aresubjected to gamma radiation sterilization nominally at 2.5-3.5 Mrad.The stents enjoy full resiliency after radiation whether exposed in aconstrained or non-constrained status. It has been found thatconstrained stents subjected to gamma sterilization without utilizingthe argon plasma pretreatment lose resiliency and do not recover at asufficient or appropriate rate.

The elastomeric material that forms a major constituent of the stentcoating should possess certain properties. It is preferably a suitablehydrophobic biostable elastomeric material which does not degrade andwhich minimizes tissue rejection and tissue inflammation and one whichwill undergo encapsulation by tissue adjacent to the stent implantationsite. Polymers suitable for such coatings include silicones (e.g.,polysiloxanes and substituted polysiloxanes), polyurethanes (includingpolycarbonate urethanes), thermoplastic elastomers in general, ethylenevinyl acetate copolymers, polyolefin elastomers, EPDM(ethylene-propylene terpolymer) rubbers and polyamide elastomers. Theabove-referenced materials are considered hydrophobic with respect tothe contemplated environment of the invention.

Agents suitable for incorporation include antithrobotics,anticoagulants, antiplatelet agents, thrombolytics, antiproliferatives,antinflammatories, agents that inhibit hyperplasia and in particularrestenosis, smooth muscle cell inhibitors, antibiotics growth factors,growth factor inhibitors, cell adhesion inhibitors, cell adhesionpromoters and drugs that may enhance the formation of healthy neointimaltissue, including endothelial cell regeneration. The positive action maycome from inhibiting particular cells (e.g., smooth muscle cells) ortissue formation (e.g., fibromuscular tissue) while encouragingdifferent cell migration (e.g., endothelium) and tissue formation(neointimal tissue).

The preferred materials for fabricating the braided stent includestainless steel, tantalum, titanium alloys including nitinol (a nickeltitanium, thermomemoried alloy material), and certain cobalt alloysincluding cobalt-chromium-nickel alloys such as ELGILOY® and PHYNOX.Further details concerning the fabrication and details of other aspectsof the stents themselves, may be gleaned from the above referenced U.S.Pat. Nos. 4,655,771 and 4,954,126 to Wallsten and 5,061,275 to Wallstenet al. To the extent additional information contained in theabove-referenced patents is necessary for an understanding of thepresent invention, they are deemed incorporated by reference herein.

Various combinations of polymer coating materials can be coordinatedwith biologically active species of interest to produce desired effectswhen coated on stents to be implanted in accordance with the invention.Loadings of therapeutic materials may vary. The mechanism ofincorporation of the biologically active species into the surfacecoating, and egress mechanism depend both on the nature of the surfacecoating polymer and the material to be incorporated. The mechanism ofrelease also depends on the mode of incorporation. The material mayelute via interparticle paths or be administered via transport ordiffusion through the encapsulating material itself.

For the purposes of this specification, "elution" is defined as anyprocess of release that involves extraction or release by direct contactof the material with bodily fluids through the interparticle pathsconnected with the exterior of the coating. "Transport" or "diffusion"are defined to include a mechanism of release in which a materialreleased traverses through another material.

The desired release rate profile can be tailored by varying the coatingthickness, the radial distribution (layer to layer) of bioactivematerials, the mixing method, the amount of bioactive material, thecombination of different matrix polymer materials at different layers,and the crosslink density of the polymeric material. The crosslinkdensity is related to the amount of crosslinking which takes place andalso the relative tightness of the matrix created by the particularcrosslinking agent used. This, during the curing process, determines theamount of crosslinking and so the crosslink density of the polymermaterial. For bioactive materials released from the crosslinked matrix,such as heparin, a crosslink structure of greater density will increaserelease time and reduce burst effect.

Additionally, with eluting materials such as heparin, release kinetics,particularly initial drug release rate, can be affected by varying theaverage dispersed particle size. The observed initial release rate orburst effect may be substantially reduced by using smaller particles,particularly if the particle size is controlled to be less than about 15microns and the effect is even more significant in the particle sizerange of ≦10 microns, especially when the coating thickness is not morethan about 50 μm and drug loading is about 25-45 weight percent.

It will also be appreciated that an unmedicated silicone thin top layerprovides an advantage over drug containing top coat. Its surface has alimited porosity and is generally smooth, which may be lessthrombogeneous and may reduce the chance to develop calcification, whichoccurs most often on the porous surface.

BRIEF DESCRIPTION OF THE DRAWINGS

In the drawings, wherein like numerals designate like parts throughoutthe same:

FIG. 1 is a schematic flow diagram illustrating the steps of the processof the invention;

FIG. 2 represents a release profile for a multi-layer system showing thepercentage of heparin released over a two-week period;

FIG. 3 represents a release profile for a multi-layer system showing therelative release rate of heparin over a two-week period;

FIG. 4 illustrates a profile of release kinetics for different drugloadings at similar coating thicknesses illustrating the release ofheparin over a two-week period;

FIG. 5 illustrates drug elution kinetics at a given loading of heparinover a two-week period at different coating thicknesses;

FIG. 6 illustrates the release kinetics in a coating having a giventie-layer thickness for different top coat thicknesses in which thepercentage heparin in the tie coat and top coats are kept constant;

FIG. 7 illustrates the release kinetics of several coatings having anaverage coating thickness of 25 microns and a heparin loading of 37.5%but using four different average particle sizes;

FIGS. 8-11 are photomicrographs of coated stent fragments for thecoatings of FIG. 7 having a corresponding average particle size of 4microns, 17 microns, 22 microns and 30 microns, respectively.

DETAILED DESCRIPTION

According to the present invention, the stent coatings incorporatingbiologically active materials for timed delivery in situ in a body lumenof interest are preferably sprayed in many thin layers from preparedcoating solutions or suspensions. The steps of the process areillustrated generally in FIG. 1. The coating solutions or suspensionsare prepared at 10 as will be described later. The desired amount ofcrosslinking agent is added to the suspension/solution as at 12 andmaterial is then agitated or stirred to produce a homogenous coatingcomposition at 14 which is thereafter transferred to an applicationcontainer or device which may be a container for spray painting at 16.Typical exemplary preparations of coating solutions that were used forheparin and dexamethasone appear next.

General Preparation of Heparin Coating Composition

Silicone was obtained as a polymer precursor in solvent (xylene)mixture. For example, a 35% solid silicone weight content in xylene wasprocured from Applied Silicone, Part #40,000. First, the silicone-xylenemixture was weighed. The solid silicone content was determined accordingto the vendor's analysis. Precalculated amounts of finely dividedheparin (2-6 microns) were added into the silicone, then tetrahydrofuron(THF) HPCL grade (Aldrich or EM) was added. For a 37.5% heparin coating,for example: W_(silicone) =5 g; solid percent =35%; W_(hep)=5×0.35×0.375/(0.625)=1.05 g. The amount of THF needed (44 ml) in thecoating solution was calculated by using the equation W_(silicone) solid/V_(THF) =0.04 for a 37.5% heparin coating solution). Finally, themanufacturer crosslinker solution was added by using Pasteur P-pipet.The amount of crosslinker added was formed to effect the release rateprofile. Typically, five drops of crosslinker solution were added foreach five grams of silicone-xylene mixture. The crosslinker may be anysuitable and compatible agent including platinum and peroxide basedmaterials. The solution was stirred by using the stirring rod until thesuspension was homogenous and milk-like. The coating solution was thentransferred into a paint jar in condition for application by air brush.

General Preparation of Dexamethasone Coating Composition

Silicone (35% solution as above) was weighed into a beaker on a Metlerbalance. The weight of dexamethasone free alcohol or acetate form wascalculated by silicone weight multiplied by 0.35 and the desiredpercentage of dexamethasone (1 to 40%) and the required amount was thenweighed. Example: W_(silicone) =5 g; for a 10% dexamethasone coating,W_(dex) =5×0.35×0.1/0.9=0.194 g and THF needed in the coating solutioncalculated. W_(silicone) solid /V_(THF) =0.06 for a 10% dexamethasonecoating solution. Example: W_(silicone) =5 g; V_(THF) =5×0.35/0.06 ≈29ml. The dexamethasone was weighed in a beaker on an analytical balanceand half the total amount of THF was added. The solution was stirredwell to ensure full dissolution of the dexamethasone. The stirredDEX-THF solution was then transferred to the silicone container. Thebeaker was washed with the remaining THF and this was transferred to thesilicone container. The crosslinker was added by using a Pasteur pipet.Typically, five drops of crosslinker were used for five grams ofsilicone.

The application of the coating material to the stent was quite similarfor all of the materials and the same for the heparin and dexamethasonesuspensions prepared as in the above Examples. The suspension to beapplied was transferred to an application device, typically a paint jarattached to an air brush, such as a Badger Model 150, supplied with asource of pressurized air through a regulator (Norgren, 0-160 psi). Oncethe brush hose was attached to the source of compressed air downstreamof the regulator, the air was applied. The pressure was adjusted toapproximately 15-25 psi and the nozzle condition checked by depressingthe trigger.

Any appropriate method can be used to secure the stent for spraying androtating fixtures were utilized successfully in the laboratory. Bothends of the relaxed stent were fastened to the fixture by two resilientretainers, commonly alligator clips, with the distance between the clipsadjusted so that the stent remained in a relaxed, unstretched condition.The rotor was then energized and the spin speed adjusted to the desiredcoating speed, nominally about 40 rpm.

With the stent rotating in a substantially horizontal plane, the spraynozzle was adjusted so that the distance from the nozzle to the stentwas about 2-4 inches and the composition was sprayed substantiallyhorizontally with the brush being directed along the stent from thedistal end of the stent to the proximal end and then from the proximalend to the distal end in a sweeping motion at a speed such that onespray cycle occurred in about three stent rotations. Typically a pauseof less than one minute, normally about one-half minute, elapsed betweenlayers. Of course, the number of coating Layers did and will vary withthe particular application. For example, for a coating level of 3-4 mgof heparin per cm² of projected area, 20 cycles of coating applicationare required and about 30 ml of solution will be consumed for a 3.5 mmdiameter by 14.5 cm long stent.

The rotation speed of the motor, of course, can be adjusted as can theviscosity of the composition and the flow rate of the spray nozzle asdesired to modify the layered structure. Generally, with the abovemixes, the best results have been obtained at rotational speeds in therange of 30-50 rpm and with a spray nozzle flow rate in the range of4-10 ml of coating composition per minute, depending on the stent size.It is contemplated that a more sophisticated, computer-controlledcoating apparatus will successfully automate the process demonstrated asfeasible in the laboratory.

Several applied layers make up what is called the tie layer as at 18 andthereafter additional upper layers, which may be of a differentcomposition with respect to bioactive material, the matrix polymericmaterials and crosslinking agent, for example, are applied as the toplayer as at 20. The application of the top layer follows the samecoating procedure as the tie layer with the number and thickness oflayers being optional. Of course, the thickness of any layer can beadjusted by modifying the speed of rotation of the stent and thespraying conditions. Generally, the total coating thickness iscontrolled by the number of spraying cycles or thin coats which make upthe total coat.

As shown at 22 in FIG. 1, the coated stent is thereafter subjected to acuring step in which the pre-polymer and crosslinking agents cooperateto produce a cured polymer matrix containing the biologically activespecies. The curing process involves evaporation of the solvent xylene,THF, etc. and the curing and crosslinking of the polymer. Certainsilicone materials can be cured at relatively low temperatures, (i.e.RT-50° C.) in what is known as a room temperature vulcanization (RTV)process. More typically, however, the curing process involves highertemperature curing materials and the coated stents are put into an ovenat approximately 90° C. or higher for approximately 16 hours. Thetemperature may be raised to as high as 150° C. for dexamethasonecontaining coated stents. Of course, the time and temperature may varywith particular silicones, crosslinkers, and biologically activespecies.

Stents coated and cured in the manner described need to be sterilizedprior to packaging for future implantation. For sterilization, gammaradiation is a preferred method particularly for heparin containingcoatings; however, it has been found that stents coated and curedaccording to the process of the invention subjected to gammasterilization may be too slow to recover their original posture whendelivered to a vascular or other lumen site using a catheter unless apretreatment step as at 24 is first applied to the coated, cured stent.

The pretreatment step involves an argon plasma treatment of the coated,cured stents in the unconstrained configuration. In accordance with thisprocedure, the stents are placed in a chamber of a plasma surfacetreatment system such as a Plasma Science 350 (Himont/Plasma Science,Foster City, Calif.). The system is equipped with a reactor chamber andRF solid-state generator operating at 13.56 mHz and from 0-500 wattspower output and being equipped with a microprocessor controlled systemand a complete vacuum pump package. The reaction chamber contains anunimpeded work volume of 16.75 inches (42.55 cm) by 13.5 inches (34.3cm) by 17.5 inches (44.45 cm) in depth.

In the plasma process, unconstrained coated stents are placed in areactor chamber and the system is purged with nitrogen and a vacuumapplied to 20-50 mTorr. Thereafter, inert gas (argon, helium or mixtureof them) is admitted to the reaction chamber for the plasma treatment. Ahighly preferred method of operation consists of using argon gas,operating at a power range from 200 to 400 watts, a flow rate of 150-650standard ml per minute, which is equivalent to 100-450 mTorr, and anexposure time from 30 seconds to about 5 minutes. The stents can beremoved immediately after the plasma treatment or remain in the argonatmosphere for an additional period of time, typically five minutes.

After this, as shown at 26, the stents are exposed to gammasterilization at 2.5-3.5 Mrad. The radiation may be carried out with thestent in either the radially non-constrained status--or in the radiallyconstrained status.

With respect to the anticoagulant material heparin, the percentage inthe tie layer is nominally from about 20-50% and that of the top layerfrom about 0-30% active material. The coating thickness ratio of the toplayer to the tie layer varies from about 1:10 to 1:2 and is preferablyin the range of from about 1:6 to 1:3.

Suppressing the burst effect also enables a reduction in the drugloading or in other words, allows a reduction in the coating thickness,since the physician will give a bolus injection ofantiplatelet/anticoagulation drugs to the patient during the stentingprocess. As a result, the drug imbedded in the stent can be fully usedwithout waste. Tailoring the first day release, but maximizing secondday and third day release at the thinnest possible coating configurationwill reduce the acute or subcute thrombosis.

FIG. 4 depicts the general effect of drug loading for coatings ofsimilar thickness. The initial elution rate increases with the drugloading as shown in FIG. 5. The release rate also increases with thethickness of the coating at the same loading but tends to be inverselyproportional to the thickness of the top layer as shown by the same drugloading and similar tie-coat thickness in FIG. 6.

The effect of average particle size is depicted in the FIGS. 7-11 inwhich coating layers with an average coating thickness of about 25microns (μm), prepared and sterilized as above, were provided withdispersed heparin particles (to 37.5% heparin) of several differentaverage particle sizes. FIG. 7 shows plots of elution kinetics for fourdifferent sizes of embedded heparin particles. The release took place inphosphate buffer (pH 7.4) at 37° C. The release rate using smaller,particularly 4-6μm average sized particles noticeably reduces theinitial rate or burst effect and thereafter the elution rate decreasesmore slowly with time. Average particle sizes above about 15 μm resultin initial release rates approaching bolus elution. This, of course, isless desirable, both from the standpoint of being an unnecessary initialexcess and for prematurely depleting the coating of deserved drugmaterial.

In addition, as shown in the photomicrographs of FIGS. 8-11, as theaverage particle size increases, the morphology of the coating surfacealso changes. Coatings containing larger particles (FIGS. 9-11) havevery rough and irregular surface characteristics. These surfaceirregularities may be more thrombogenic or exhibit an increased tendencyto cause embolization when the corresponding stent is implanted in ablood vessel.

Accordingly, it has been found that the average particle size shouldgenerally be controlled below about 15 μm to reduce the burst effect andpreferably should be ≦about 10 μm for best results. The 4-6 μm sizeworked quite successfully in the laboratory. However, it should be notedthat larger particle size can also be advantageously used, for instance,when the drug load is low, such as below 25 weight percent. Elutionkinetics can be adjusted by a combination of changing the particle sizeand changing the load or concentration of the dispersed drug material.

What is apparent from the data gathered to date, however, is that theprocess of the present invention enables the drug elution kinetics to bemodified to meet the needs of the particular stent application. In asimilar manner, stent coatings can be prepared using a combination oftwo or more drugs and the drug release sequence and rate controlled. Forexample, antiproliferation drugs may be combined in the undercoat andanti-thrombotic drugs in the topcoat layer. In this manner, theanti-thrombotic drugs, for example, heparin, will elute first followedby antiproliferation drugs, e.g. dexamethasone, to better enable safeencapsulation of the implanted stent.

The heparin concentration measurement were made utilizing a standardcurve prepared by complexing azure A dye with dilute solutions ofheparin. Sixteen standards were used to compile the standard curve in awell-known manner.

For the elution test, the stents were immersed in a phosphate buffersolution at pH 7.4 in an incubator at approximately 37° C. Periodicsamplings of the solution were processed to determine the amount ofheparin eluted. After each sampling, each stent was placed inheparin-free buffer solution.

As stated above, while the allowable loading of the elastomeric materialwith heparin may vary, in the case of silicone materials heparin mayexceed 60% of the total weight of the layer. However, the loadinggenerally most advantageously used is in the range from about 10% to 45%of the total weight of the layer. In the case of dexamethasone, theloading may be as high as 50% or more of the total weight of the layerbut is preferably in the range of about 0.4% to 45%.

It will be appreciated that the mechanism of incorporation of thebiologically active species into a thin surface coating structureapplicable to a metal stent is an important aspect of the presentinvention. The need for relatively thick-walled polymer elution stentsor any membrane overlayers associated with many prior drug elutiondevices is obviated, as is the need for utilizing biodegradable orreabsorbable vehicles for carrying the biologically active species. Thetechnique clearly enables long-term delivery and minimizes interferencewith the independent mechanical or therapeutic benefits of the stentitself.

Coating materials are designed with a particular coating technique,coating/drug combination and drug infusion mechanism in mind.Consideration of the particular form and mechanism of release of thebiologically active species in the coating allow the technique toproduce superior results. In this manner, delivery of the biologicallyactive species from the coating structure can be tailored to accommodatea variety of applications.

Whereas the above examples depict coatings having two different drugloadings or percentages of biologically active material to be released,this is by no means limiting with respect to the invention and it iscontemplated that any number of layers and combinations of loadings canbe employed to achieve a desired release profile. For example, gradualgrading and change in the loading of the layers can be utilized inwhich, for example, higher loadings are used in the inner layers. Alsolayers can be used which have no drug loadings at all. For example, apulsatile heparin release system may be achieved by a coating in whichalternate layers containing heparin are sandwiched between unloadedlayers of silicone or other materials for a portion of the coating. Inother words, the invention allows untold numbers of combinations whichresult in a great deal of flexibility with respect to controlling therelease of biologically active materials with regard to an implantedstent. Each applied layer is typically from approximately 0.5 microns to15 microns in thickness. The total number of sprayed layers, of course,can vary widely, from less than 10 to more than 50 layers; commonly, 20to 40 layers are included. The total thickness; of the coating can alsovary widely, but can generally be from about 10 to 200 microns.

Whereas the polymer of the coating may be any compatible biostableelastomeric material capable of being adhered to the stent material as athin layer, hydrophobic materials are preferred because it has beenfound that the release of the biologically active species can generallybe more predictably controlled with such materials. Preferred materialsinclude silicone rubber elastomers and biostable polyurethanesspecifically.

This invention has been described herein in considerable detail in orderto comply with the Patent Statutes and to provide those skilled in theart with the information needed to apply the novel principles and toconstruct and use embodiments of the example as required. However, it isto be understood that the invention can be carried out by specificallydifferent devices and that various modifications can be accomplishedwithout departing from the scope of the invention itself.

We claim:
 1. A method of coating at least a portion of an implantableprosthesis, having at least one opening therein, with a hydrophobicelastomeric material incorporating an amount of biologically activematerial therein for timed delivery therefrom comprising the stepsof:(a) applying a coating comprising the elastomeric material, a solventand an amount of finely divided biologically active material onto atleast a portion of the prosthesis; wherein when the biologically activematerial is particulate the average particle size of the biologicallyactive material is less than or equal to about 15 μm; and wherein thecoating is applied to the prosthesis in a manner to adheringly conformthereto to preserve the opening; and (b) curing the coating such that atleast some of the biologically active material is particulate aftercuring.
 2. The method of claim 1 wherein the elastomeric material isselected from the group consisting of silicones, polyurethanes,polyamide elastomers, ethylene vinyl acetate copolymers, polyolefinelastomers, ethylene-propylene terpolymer rubbers and combinationsthereof.
 3. The method of claim 1 wherein the biologically activematerial includes heparin.
 4. The method of claim 1 wherein the coatingcomprises about 25-45 weight percent biologically active material. 5.The method of claim 1 wherein the biologically active material has anaverage particle size less than or equal to about 10 μm before curing.6. The method of claim 5 wherein the biologically active materialincludes heparin.
 7. A method of controlling the delivery of an elutingmaterial incorporated in an elastomeric coating having at least onelayer on at least a portion of an implantable prosthesis having at leastone opening therein, the method comprising incorporating a biologicallyactive particulate material having an average particle size of less thanor equal to about 15 μm into at least one layer of the coating andapplying the elastomeric coating in a manner which adheringly conformsto the surface to preserve the opening; and curing the coating such thatat least some of the biologically active material is particulate aftercuring.
 8. The method of claim 7 wherein said biologically activematerial is heparin.
 9. The method of claim 7 wherein the layercomprises about 25-45 weight percent biologically active material. 10.The method of claim 7 wherein the biologically active material isincorporated to produce a substantially smooth surface on theprosthesis.
 11. The method of claim 1 wherein the elastomeric material,solvent and biologically active material are applied by spraycoating theprosthesis.
 12. The method of claim 1 wherein the elastomeric material,solvent and biologically active material are applied by dipping theprosthesis.
 13. The method of claim 7 wherein the biologically activematerial has an average particle size less than or equal to about 10 μmbefore curing.
 14. The method of claim 1 wherein the implantableprosthesis is an expandable stent having a tubular metal body havingopen ends and a sidewall structure having openings therein, and whereinthe elastomeric material, solvent and biologically active material forma coating on a surface of said sidewall structure which continuouslyconforms to said sidewall structure in a manner that preserves theopenings when the stent is expanded.
 15. The method of claim 13 whereinthe elastomeric material, solvent and biologically active material areapplied with the stent fully expanded.
 16. The method of claim 1 whereinthe elastomeric material, solvent and biologically active material arein a mixture.
 17. The method of claim 1 wherein the biologically activematerial has an average particle size of less than or equal to about 15μm after curing.
 18. The method of claim 5 wherein the biologicallyactive material has an average particle size of less than or equal toabout 15 μm after curing.